1. Field of the Invention
This invention relates to a rotary-to-linear motion conversion device suitable for use where redundant rotary actuation is required. In particular, the invention relates to a device for actuating a pneumatic pump driving an artificial heart.
2. Description of the Prior Art
The heart is the muscle that drives the cardiovascular system in living beings. Acting as a pump, the heart moves blood throughout the body to provide oxygen, nutrients, hormones, and to remove waste products. The blood follows two separate pathways in the human body, the so-called pulmonary and systemic circulatory circuits. In the pulmonary circuit, the heart pumps blood first to the lungs to release carbon dioxide and bind oxygen, and then back to the heart. Thus, oxygenated blood is constantly being supplied to the heart. In the systemic circuit, the longer of the two, the heart pumps oxygenated blood through the rest of the body to supply oxygen and remove carbon dioxide, the byproduct of metabolic functions carried out throughout the body. The heart supplies blood to the two circuits with pulses generated by the orderly muscular contraction of its walls.
In order to keep blood moving through these two separate circulatory circuits, the human heart has four distinct chambers that work in pairs. As illustrated in FIG. 1, the heart 10 includes a right atrium 12, a right ventricle 14, a left atrium 16, and a left ventricle 18. One pair of chambers, the right ventricle and left atrium, is connected directly to the pulmonary circuit. In it, de-oxygenated blood from the body is pumped from the right ventricle 14 to the lungs, where it is oxygenated, and then back to the left atrium 16.
In the systemic circuit, the other pair of chambers pumps the oxygenated blood through body organs, tissues and bones. The blood moves from the left atrium 16, where it flows from the lungs, to the left ventricle 18, which in turn pumps the blood throughout the body and all the way back to the right atrium 12. The blood then moves to the right ventricle 14 where the cycle is repeated. In each circuit, the blood enters the heart through an atrium and leaves the heart through a ventricle.
Thus, the ventricles 14,18 are essentially two separate pumps that work together to move the blood through the two circulatory circuits. Four check valves control the flow of blood within the heart and prevent flow in the wrong direction. A tricuspid valve 20 controls the blood flowing from the right atrium 12 into the right ventricle 14. Similarly, a bicuspid valve 22 controls the blood flowing from the left atrium 16 into the left ventricle 18. Two semilunar valves (pulmonary 24 and aortic 26) control the blood flow leaving the heart toward the pulmonary and systemic circuits, respectively. Thus, in each complete cycle, the blood is pumped by the right ventricle 14 through the pulmonary semilunar valve 24 to the lungs and back to the left atrium 16. The blood then flows through the bicuspid valve 22 to the left ventricle 18, which in turn pumps it through the aortic semilunar valve 26 throughout the body and back to the right atrium 12. Finally, the blood flows back to the right ventricle 14 through the tricuspid valve 20 and the cycle is repeated.
When the heart muscle squeezes each ventricle, it acts as a pump that exerts pressure on the blood, thereby pushing it out of the heart and through the body. The blood pressure, an indicator of heart function, is measured when the heart muscle contracts as well as when it relaxes. The so-called systolic pressure is the maximum pressure exerted by the blood on the arterial walls when the left ventricle of the heart contracts forcing blood through the arteries in the systemic circulatory circuit. The so-called diastolic pressure is the lowest pressure on the blood vessel walls when the left ventricle relaxes and refills with blood. Healthy blood pressure is considered to be about 120 millimeters of mercury systolic and 80 millimeters of mercury diastolic (usually presented as 120/80).
Inasmuch as the function of the circulatory system is to service the biological needs of all body tissues (i.e., to transport nutrients to the tissues, transport waste products away, distribute hormones from one part of the body to another, and, in general, to maintain an appropriate environment for optimal function and survival of tissue cells), the rate at which blood is circulated by the heart is a critical aspect of its function. The heart has a built-in mechanism (the so-called Frank-Starling mechanism) that allows it to pump automatically whatever amount of blood flows into it. Such cardiac output in a healthy human body may vary from about 4 to about 15 liters per minute (LPM), according to the activity being undertaken by the person, at a heart rate that can vary from about 50 to about 180 beats per minute.
Several artificial devices have been developed over the years to supplement or replace the function of a failing heart in patients. These include devices developed by companies as well as research institutions such as the Berlin Heart Institute, the Pennsylvania State University, the University of Utah, the Cleveland Clinic Foundation, the University of Perkinje (in Bruno, Czechoslovakia), the University of Tokyo, the Thoratec Corporation, Abiomed Inc., Novacor, and Symbion Inc. Typically, these artificial devices consist of pumps that aim at duplicating the required pumping functions of the left and right human ventricles. One method of actuation for these pumps has been through the pneumatic action of an external mechanism. See, for example, U.S. Pat. Nos. 4,611,578 and 5,766,207. Periodic pulses of compressed air drive the pumps at the desired pressure and rate of cardiac output. A moderate vacuum may be applied between pulses to allow more rapid refilling of the ventricles with blood flowing from the respective atrium.
One notable artificial heart currently in use as an implant for patients waiting for a heart transplant is the Total Artificial Heart manufactured by SynCardia Systems, Inc., of Tucson, Ariz. Designed to operate much the same way as a human heart, this artificial heart replaces the two active chambers (i.e., the ventricles) of the human heart with corresponding artificial components. As illustrated in FIG. 2, such artificial heart 30 includes two separate chambers or ventricles 32 and 34 that replace the right and left ventricles of the human heart, respectively. Each chamber is equipped with a respective diaphragm (36 and 38 in the right and left chamber, respectively) that has an air contact side and a blood contact side. Each diaphragm is designed as a spherical hemisphere. As shown in FIG. 3, the artificial heart 30 is implanted by connecting the top of the right chamber 32 to the right atrium 12 and the top of the left chamber 34 to the left atrium 16. The bottom of each chamber is provided with an air line (40 and 42 in the right and left chamber, respectively) that is embedded in the patient's body but extends outside for connection to a pneumatic driver.
When driven by a supply of pressurized air from the pneumatic driver, each diaphragm 36,38 discharges blood from the respective chamber 32,34 simulating the function of a ventricle. This phase is referred to in the art as systole or equivalently as the ejection phase. When the pressurized air is removed from the diaphragm, known as diastole or the filling phase, blood can enter the ventricle from the connected atrium. The rate at which blood enters the ventricle depends on the difference between the atrial pressure and the pressure on the air-side of the diaphragm. To increase this filling rate, a slight vacuum of about 10 mm Hg is normally applied to the air-side of the diaphragm during diastole. Artificial valves 44a (tricuspid), 46a (bicuspid) and 44b (pulmonary), 46b (aortic) control the flow from the respective atrium into each artificial ventricle and out to the circulatory systems, respectively.
The pneumatic drivers used to date for driving all artificial hearts have been cumbersome and inadequate for affording patients any degree of independent mobility. They employ compressors, vacuum pumps, and air tanks coupled to electrically actuated valves, all of which amounts to a large and heavy apparatus that can only be transported on wheels and with considerable effort. Therefore, many attempts have been made during the last two decades to produce a portable driver for these devices. However, because of the complexity of the required functionality and the hardware necessary to produce it, pneumatic heart drivers continue to be bulky, require frequent maintenance, and often provide air pulses that do not match the performance of the larger drivers they are meant to replace. Even at the approximate weight of 20 pounds and size of about 0.7 cubic feet achieved so far, pneumatic drivers remain unwieldy and substantially not portable for a patient who is kept alive by an artificial heart.
In essence, a portable driver needs to be reliable, durable, easy to use, and sufficiently simple in design to be affordable. Unfortunately, each of these requirements contributes to the complexity of the design, which in turn has produced devices that are not sufficiently small and light-weight to be manageable in the hands of a patient. Furthermore, it is essential that the pneumatic driver be able to provide the correct pressure balance between the left and right ventricles of the artificial heart to ensure the proper operating pressure to the pulmonary and systemic circuits regardless of the speed of operation. Typically, this requires that the driver be able to operate so as maintain, on average, a right atrial pressure of about 9 mmHg, a mean pulmonary artery pressure of about 35 mmHg, a left atrial pressure of about 10 mmHg, and a mean aortic pressure of about 95 mmHg.
This need to provide different operating pressures to the right and left chambers (ventricles) of the artificial-heart device has not been met heretofore with a simple design suitable for a portable driver. For example, the blood pump described in U.S. Pat. No. 4,611,578 includes a configuration wherein two reciprocating pistons in a common cylinder may be operated alternatively to provide redundancy or independently to actuate two separate pneumatically driven blood pumps. This issue is not addressed in the patent, but it describes a sophisticated control system that arguably could be used to provide the correct operating pressure to each chamber of the artificial heart. However, the complex and multi-component structure of the device necessarily requires a relatively heavy and large apparatus, though described as portable. The commercially available module weighs about 25 pounds and is approximately 0.6 cubic feet in volume.
U.S. Pat. No. 5,766,207 describes another portable pneumatic driver for ventricular assist devices that could also be adapted for an artificial heart. The single pump of the invention could be used to drive both ventricles of an artificial heart, but only at the same pressure and volume rate. Thus, this device, even if modified to meet the other requirements of a portable artificial-heart driver, would not be suitable as an alternative to the stationary modules currently in use.
Copending U.S. application Ser. No. 12/454,440, hereby incorporated by reference, describes a portable driver configured to optimize size, weight, reliability, durability, extended battery life, ease of use, and simplicity of design. The driver consists of a pneumatic pump that provides coordinated and independent periodic actuation pressure to each ventricle of the artificial heart, limiting peak pressures and peak vacuums to provide a safe and efficient cycle of operation, allowing only partial filling of each ventricle of the cardiac device to ensure redundancy of capacity, providing sufficient pneumatic stroke to completely eject the blood from the ventricles at each beat, readily adjusting the rate at which the artificial heart is actuated, and minimizing overall size and weight to enable portability.
In the preferred embodiment, the pneumatic pump 50 comprises two coaxial cylindrical pumping chambers (60 and 62) defined by a common housing 52, each enclosing a disk-shaped piston (54 and 56) incorporating seals 66 to eliminate leakage, as illustrated schematically in FIG. 4. The pistons are connected to one another through a partition 58 by a tube 64, thereby forming a monolithic piston assembly that is driven axially by a common electrical actuator 68 providing reciprocating motion through a rod connected to the top piston. The tube travels through a seal in the partition that separates the two chambers and, by defining the boundary between the pistons, also acts as a bulkhead for the top chamber.
The volume in the bottom chamber is selected as needed to provide the desired pressure in the left ventricle of the artificial heart driven by the pump. According to one aspect of the invention, the diameter of the tube connecting the pistons is selected such that the stroke volume (i.e., the displacement) of the top chamber is reduced with respect to that of the bottom chamber as needed to match the reduced pressure requirements of the right ventricle of the artificial heart. Namely, the maximum pressure achieved at the respective output port (70 and 72) in each chamber should be as needed to fully eject blood from each ventricle of the artificial heart substantially at the operating pressures of the human pulmonary and systemic circulatory circuits. A limit check valve (74 and 76) is preferably used in each chamber to ensure venting of excess pressure during the compression stroke. A limit check valve (78 and 80) is also preferably used in each chamber to limit the vacuum generated during the reverse, aspiration stroke.
FIG. 5 illustrates in sectioned view the actual pneumatic driver that incorporates the concepts of the invention disclosed in Ser. No. 12/454,440. As clearly illustrated by FIGS. 4 and 5, the pneumatic pump is actuated by a reciprocating mechanism that drives the pistons. Many such mechanisms are known in the art that could be used to operate the pump. The device commonly referred to as a scotch yoke is preferred because of its simplicity, reliability and suitability for implementation in a small volume, all of which are critical for a portable artificial heart driver.
However, an additional requirement for such an application is redundancy, which is not readily available from a conventional application of scotch-yoke technology wherein a motor 68 (FIG. 5) with a rotating output shaft is the driving source for actuating the reciprocating pistons of the pump. The present invention provides an ingenious solution to this problem in a configuration designed particularly to meet the critical redundancy requirements of the SynCardia artificial heart driver.